1. Field of the Invention
The present invention relates generally to pumps which employ magnetic suspension and rotation means to pump blood, and more particularly to a magnetically suspended and rotated blood pump that has no mechanical bearings or seals and has a pump means which is magnetically supported radially and axially.
2. Description of the Prior Art
The use of rotary pumps (i.e. axial, centrifugal, mixed flow) to pump fluids and in particular blood is well known by those skilled in the art. A rotary pump, in general, consists of an outer housing, with inlet and outlet ports, and an impeller mounted on a shaft (with mechanical bearings and seals) within the outer housing for rotation about an axis. Mechanical bearings are susceptible to wear and premature failure and can generate sufficient heat and mechanical stresses to cause unacceptable blood damage. Shaft seals are also susceptible to wear and heat generation, which can lead to leakage, blood clot formation, bearing seizure, and bacterial growth. Examples of rotary pumps utilizing shaft mounted impellers with bearings and seals are disclosed in Reich et. al. U.S. Pat. No. 4,135,253; Possell U.S. Pat. No. 4,403,911; Moise U.S. Pat. No. 4,704,121; and Dorman U.S. Pat. No. 4,927,407.
Numerous pumps have been designed to circumvent the above problems by employing a lubricant flush of rotary pump mechanical bearings. Examples of such pumps are disclosed in Carriker et al. U.S. Pat. No. 4,944,722 and Wampler et al. U.S. Pat. No. 4,846,152. These types of pumps can have several problems including not having the ability to be fully implantable due to the need for a percutaneous supply line and external reservoir to achieve bearing flushing. Also the potential for infection and leakage exists due to the flushing fluid and percutaneous lines. In addition the mechanical bearings can still require replacement after time because they directly contact other pump structures during operation.
By employing a rotary fluid pump with a magnetically suspended impeller, all of the above mentioned problems can be avoided. Examples of such pumps are disclosed in Bramm et al. U.S. Pat. No. 5,326,344; Olsen et al. U.S. Pat. No. 4,688,998 and Moise U.S. Pat. No. 4,779,614. A problem which can be associated with all of the cited inventions is that a single gap is employed for both the blood flow pathway through the pump and for the magnetic suspension and rotation of the impeller. These two functions have directly opposing requirements on the size of the gap. As a blood flow pathway, the gap should be large to avoid blood damage. As a magnetic suspension and rotation gap, the gap should be small to minimize the size of the magnetic suspension and rotation components and also to allow for efficient use of energy to achieve impeller suspension and rotation. Consequently, for these types of pumps, any gap size selected can result in an undesirable compromise between blood damage, device size, and energy requirements.
Examples of pumps having separate gaps for primary blood flow and impeller rotation are disclosed in Golding et al. U.S. Pat. No. 5,324,177 and Golding et al. U.S. Pat. No. 5,049,134. However, these pumps also use the rotation gap to implement hydrodynamic suspension bearings for the rotor. Such hydrodynamic bearings can subject the blood to excessive shear stresses which can unacceptably damage the fragile components of the blood. Additionally, the Golding et. al. pumps place the stationary magnetic components inside a center-bore of a rotating assembly. Such configurations generally cause the mass and rotational inertia of the rotating assembly to be larger than those in a system in which the stationary magnetic components are placed around the outer surface of the rotating assembly. Rotating assemblies having large masses and rotational inertias can be undesirable because the axial and radial bearing elements must be made relatively large in order to maintain proper alignment of the rotating assembly during shock, vibration, and acceleration.
The flow rate of blood pumps that are capable of creating negative inlet pressures must be dynamically adjusted to match the blood flow rate into the ventricle of the heart, typically the left ventricle. If too little flow is produced by the blood pump, the tissues and organs of the body may be inadequately perfused, and the blood pressure in the left ventricle will increasexe2x80x94potentially causing excessive pulmonary pressure and congestion. Conversely, if the flow rate of the blood pump is too high, excessive negative pressure may be created in the left ventricle and in the inlet to the pump. Excessive negative blood pressure is undesirable for the following reasons: 1) Unacceptable levels of blood damage may be caused by cavitation; 2) The pump may be damaged by cavitation; 3) The walls of the ventricle may collapse and be damaged; and 4) The walls of the ventricle may collapse and block the blood flow pathway to the pump.
By employing a control system to dynamically control the flow rate of the pump to avoid excessive negative blood pressure the above mentioned problems can be avoided. One example of such a control system is disclosed in Bramm et al., U.S. Pat. No. 5,326,344. Bramm describes a method of dynamically controlling the flow rate of a pump based on a signal derived from a single pressure sensor located within the pump inlet. One problem which can be associated with such a pressure sensing system is the difficulty in achieving long-term stability of such a sensor, particularly in light of the relatively low pressures (0 to 20 mm Hg) that must be resolved and the hostile environment in which the sensor is operated. Another problem which can be associated with such a pressure sensing system is that the effect of changing atmospheric pressure can cause inaccurate sensing of the pressure needed to properly control the pump.
Many patients that are in need of cardiac assistance due to their heart""s inability to provide adequate blood flow are also predisposed to cardiac arrhythmias. Such arrhythmias can adversely affect blood flow when a cardiac assist device is used, particularly when only uni-ventricular cardiac assistance is being provided. By combining an arrhythmia control system with a cardiac assistance system, the above mentioned problems can be alleviated. One example of such a combined cardiac assist and arrhythmia control system is disclosed by Heilman et al. U.S. Pat. No. 4,925,443. Heilman describes a cardiac assist device that directly compresses the myocardium to achieve increased blood flow combined with an arrhythmia control system. Some problems which can be associated with direct compression of the myocardium can include difficulty in conforming to a wide range of heart shapes and sizes, difficulty in adequately attaching such a device to the heart, and damage of the myocardium due to compression and abrasion.
Accordingly, there is a need for a blood pump which overcomes the aforementioned problems that can be associated with conventional blood pumps and also a system of dynamically controlling such a blood pump to avoid the previously described problems that can occur with control systems using pressure sensors. Moreover, such blood pump and control system should be able to cooperate with an arrhythmia control system for improved cardiac arrhythmia treatment.
A blood pump apparatus is provided which can include a stator member containing a magnetically suspended and rotated rotor member. The rotor can preferably be magnetically suspended within the stator both radially and axially. The blood pump can also have an associated magnetic suspension control system, a blood pump flow rate control system, and an arrhythmia control system. The blood pump can preferably be a centrifugal pump wherein an impeller draws blood from the left ventricle of a the heart and delivers it to the aorta thereby reducing the pressure that must be generated by the left ventricle. The blood pump can also be of a relatively small size such that it can be completely implanted within the human body. If bi-ventricular cardiac assist is needed a second such blood pump can be implanted to assist the right ventricle. The impeller of the centrifugal pump can be an integral part of a rotor assembly. The rotor assembly can preferably be suspended by permanent magnet radial bearings and a Lorentz-force axial bearing. The Lorentz-force axial bearing can generate bi-directional axial forces in response to an applied current. The blood pump can also include an axial position sensor and an axial position controller. The axial position sensor can monitor the axial position of the rotor and provide feedback to the controller to maintain the axial position of the rotor. The axial position controller can also adjust the axial position of the rotor such that steady-state axial loads due to gravity, acceleration or the centrifugal pump impeller are offset by the inherent axial forces generated by the permanent magnet radial bearings. By offsetting the steady-state axial forces using the axial position controller, the power required by the Lorentz-force axial bearing is minimized. The rotor assembly can be rotated by an electric motor.
A primary blood flow inlet path can preferably be through a relatively large center bore provided in the rotor. A secondary blood flow inlet path can be through an annular gap which is formed between the rotor and the stator of the pump as a result of the radial magnetic suspension. In order to minimize the size of the device, all of the magnetic suspension and rotation forces can be applied across the relatively small annular gap. All blood contacting surfaces of the pump are continuously washed to avoid blood clots and protein deposition.
The speed of the centrifugal pump can be dynamically controlled to avoid excessive negative pressure in the left ventricle. The blood pump flow rate control system can include an electronic heart caliper. The heart caliper can be operatively attached to the outside surface of the heart and provide feedback to the blood pump flow rate control system. The heart caliper can be utilized to monitor the outside dimension of the left ventricle. The blood pump flow rate control system can preferably operate in two modes, continuous and pulsatile. In the continuous mode of operation, the pump speed can be controlled to hold the sensed left ventricle dimension at a defined setpoint. In the pulsatile mode of operation, the pump speed can be dynamically adjusted to cause the sensed left ventricle dimension to alternate between two predefined setpoints.
The blood pump can also be utilized to improve the functioning of an arrhythmia control system. Electrodes placed in or on the surface of the heart combined with an associated arrhythmia control system can be provided to detect and treat cardiac arrhythmias including bradycardia, tachycardia, and fibrillation. In order to reduce the energy needed for the arrhythmia control system to treat fibrillation, the blood pump flow rate control system can be employed to purposely reduce the radial dimension of the ventricle prior to delivering a defibrillation pulse. By minimizing the amount of blood within the ventricle chamber (a direct result of reducing the radial dimension thereof), a larger fraction of the defibrillation energy supplied by the arrhythmia control system is delivered to the myocardium, where it is needed, and a smaller fraction of the energy is delivered to the blood, where it is unnecessary.
Other details, objects, and advantages of the invention will become apparent from the following detailed description and the accompanying drawing figures of certain presently preferred embodiments thereof.